EP3458841B1 - Laser microscope with ablation function - Google Patents

Description

The invention relates to a laser microscope with which a sample can be examined both with imaging, especially non-linear optical methods, and also specifically modified by removing material, as well as a method for operation.

State of the art

In laser microscopy, a laser beam is scanned over the surface of the sample to be examined with a scanning lens. The spatial resolution of the illuminated area on the sample is diffraction-limited. The use of non-linear optical effects increases the resolution, since a signal is only generated in the region with the highest light intensity. The diameter of this region is reduced ^{by the factor n -1/2 for nonlinear effects of order n.}

The use of non-linear effects for imaging also requires the transition from continuous to pulsed lighting. The effects depend on the light intensity as a square or with an even higher power, so that they only deliver a usable signal above a certain minimum intensity. Letting this intensity take effect over the long term would require a very high level of technical effort and at the same time destroy the sample through heating. The laser energy is therefore concentrated in short pulses with a high instantaneous intensity, with the average power deposited in the sample being selected so that the sample is not excessively heated.

Such laser microscopes are made, for example, from ( T. Meyer, M. Baumgartl, T. Gottschall, T. Pascher, A. Wuttig, C. Matthäus, BFM Romeike, BR Brehm, J. Limpert, A. Tünnermann, O. Guntinas-Lichius, B. Dietzek, M. Schmitt, J. Popp, "A compact microscope setup for multimodal nonlinear imaging in clinics and its application to disease diagnostics", Analyst 138 (14), 4048-57 (2013 )) as ( T. Meyer, M. Chemnitz, M. Baumgartl, T. Gottschall, T. Pascher, C. Matthäus, BFM Romeike, BR Brehm, J. Limpert, A. Tünnermann, M. Schmitt, B. Dietzek, J. Popp, "Expanding Multimodal Microscopy by High Spectral Resolution Coherent Anti-Stokes Raman Scattering Imaging for Clinical Disease Diagnostics", Analytical Chemistry 85, 6703-6715 (2013 )) known. These microscopes combine coherent Raman scattering, i.e. stimulated Raman scattering (SRS) and coherent anti-Stokes Raman scattering (CARS), two-photon fluorescence (two-photon excited fluorescence, TPEF) and the use of the second harmonic of the excitation light (second-harmonic generation, SHG) and, if necessary, higher harmonics. Such multimodal imaging is advantageous in two ways: On the one hand, a molecule-specific contrast can be generated with which, for example, in clinical diagnostics, pathologically altered tissue can be distinguished from healthy tissue. On the other hand, the non-linear dependence of the mentioned effects on the light intensity means that of the typically Gaussian-distributed spatial intensity profile of the pulse, only the center with the highest intensity contributes to the imaging. The spatial resolution is therefore better than would be expected on the basis of the diffraction limit.

US 2008/315 119 A1 discloses a method for photo modification of a sample, in which the sample is irradiated and one or more signals are detected based on the irradiated sample. Based on an analysis of the detected signals, the lighting parameters are adjusted so that changes can be made interactively during the experiment.

( T. Minamikawa, H. Niioka, T. Araki, M. Hashimoto, "Real-time imaging of laser-induced membrane disruption of a living cell observed with multifocus coherent anti-Stokes Raman scattering microscopy", Journal of Biomedical Optics 16 (2), 021111 (2011 )) reveals the real-time imaging of the laser-induced destruction of cell membranes in a living organism as well as the cellular response with a multifocus CARS microscope. When the cell membrane is at the focal point of a If the near-infrared pulsed laser beam is destroyed, the CARS signal contribution originating from the cell membrane disappears and a new CARS signal contribution originating from an accumulation of lipids appears.

US 2009/290 150 A1 discloses a laser microscope for simultaneous imaging using CARS and multiphoton fluorescence.

DE 10 2005 044422 A1 discloses another CARS microscope with improved suppression of the non-resonant CARS signal background.

US 2016/103 072 A1 discloses a method for observing cells which, among other things, determines which cells in the sample are still alive.

US 2010/177 307 A1 discloses an optical arrangement having an optically parametric oscillator which is pumped synchronously to generate correlated pulse pairs of a signal beam and an idle beam. The phases of these correlated pulses are locked to the phase of the pump beam.

Task and solution

It is the object of the present invention to expand the known laser microscopes by the possibility of working in the near-infrared spectral range and by the possibility of locally modifying the sample with high precision.

This object is achieved according to the invention by a laser microscope with a laser source according to the main claim and by a method for operation according to the secondary claim. Further advantageous refinements emerge from the subclaims that refer back to them.

Subject of the invention

In the context of the invention, a laser microscope with a laser source was developed. This laser microscope comprises at least one first laser source that emits at least one, in particular pulsed, excitation beam, scanning optics that are designed to raster the excitation beam over the surface of a sample, focusing optics that are designed to focus the excitation beam on the sample, and at least a detector for light which the sample emits due to an optical effect in response to the excitation beam. The laser microscope can be designed for multimodal imaging, for example.

The optical effect can be a linear effect. The imaging can then take place particularly quickly because a lot of signal intensity is available. The optical effect is particularly advantageous, however, non-linear, i.e. the detector is sensitive to light which the sample emits due to a non-linear optical effect in response to the excitation beam. Then this answer comes essentially from the central area of the beam profile of the excitation beam, in which the current intensity is at a maximum.

According to the invention, a second laser source for a pulsed ablation beam is provided for the purpose of local ablation of the material of the sample, the ablation beam being guided to the sample via the scanning optics and the focusing optics.

The laser microscope uses non-linear imaging methods such as SRS, CARS, SHG and / or TPEF.

It was recognized that this combination makes it possible to select structures from an imaged area and to ablate only these structures in a highly selective manner. The fact that a significantly greater intensity is required for ablation than for imaging does not mean that the spatial resolution during ablation is worse must than when imaging. By suitable selection of the laser parameters, the pulses of the ablation beam can be designed in such a way that they interact directly with the electron shells of the atoms of the sample material in the sample and ionize them. As a result, the sample material is locally evaporated by converting the electrons into a plasma. If the ablation pulse is short enough, it only interacts with the sample material in this way, so that in particular no heat is deposited in the sample in the form of excited states. The ionization of the electron shells requires an instantaneous intensity that is only present in the immediate center of the spatial intensity distribution of the ablation pulse. The spatial resolution of the ablation is therefore at least as good as the spatial resolution of the imaging, if not even better.

Both the excitation beam and the ablation beam have an inhomogeneous intensity distribution not only laterally, that is, in the plane perpendicular to the direction of propagation. Rather, the intensity fluctuates in each case in the direction of propagation. As a result, the center of the greatest intensity is not only localized laterally, but also strongly in the direction of propagation. The ablated volume is in the order of magnitude of single cells, approx. 1 μl (picolitre). The imaged or ablated area can therefore not only be selected laterally, but also in relation to the depth below the surface of the sample. In this way, for example, structures within a biological sample can be examined and specifically changed without first having to open the surface of the sample at the location of these structures in a destructive manner. Both the excitation and the ablation beam can penetrate the sample up to a depth of a few 100 µm, for example.

The joint guidance of the excitation beam and the ablation beam via the same scanning optics and focusing optics ensures, on the one hand, that a systematic offset between the points at which both beams arrive on the sample is minimized. On the other hand, the adjustment effort is also minimized. As a result, the successful use of the laser microscope is no longer possible assumes that the user is an expert in laser microscopy. Rather, the laser microscope, including the new ablation function, is also accessible to users who are merely experts with regard to the interpretation of the images, such as doctors or biologists when used in clinical diagnostics. In particular for such applications it is furthermore advantageous that the common use of the scanning optics and focusing optics for the excitation and for the ablation enables the integration of both functions in one compact device.

In the case of non-linear imaging, the excitation states in the sample material are changed by photons of the excitation beam. For this, the energy of the photon must match the energy difference between the excitation states. The imaging therefore requires an excitation beam with one or more specific wavelengths that are matched to the sample material and to the effect to be used for the imaging. Due to the non-linear imaging, both imaging and ablation can take place with NIR lasers, the imaging methods generating signals in the visible range. This means that the same optics can be used for both tasks.

The inventors have recognized that, in contrast to this, the nonlinear ablation is essentially independent of the wavelength of the ablation beam. The partial ionization of the electron shells of atoms in the sample material is directly caused by the momentary electric field that acts on the electrons. The oscillation frequency of this electric field, and thus the wavelength of the ablation beam, is therefore irrelevant. This wavelength can therefore be freely selected according to practical or apparatus-related considerations, in particular in the near-infrared (NIR) range, in order to enable a high depth of penetration into tissue.

The fact that the excitation beam, on the one hand, and the ablation beam, on the other hand, thus interact with the sample in a qualitatively completely different manner, causes thus that the pulses of the ablation beam must be significantly shorter than the pulses of the excitation beam in order to ensure a low average power, so that no damage occurs outside of the focus by the ablation laser. The maximum instantaneous intensity of an ablation pulse is typically about a factor of 1000 greater than the maximum instantaneous intensity of an excitation pulse. Accordingly, an ablation pulse can, for example, have a pulse energy in the range between 0.1 μJ and 10 μJ with a pulse duration of 100 fs, while an excitation pulse, for example, can only have a pulse energy in the range between 1 nJ and 10 nJ with approximately 10 ps pulse duration.

A particular advantage of integrating the laser microscope and ablation tool in one device is that the ablation can be interrupted at any time and visually checked by making a new microscope image or images of the process can be recorded permanently during the ablation. In this way there is an online control with regard to the selectivity with which the sample material is removed.

In a particularly advantageous embodiment of the invention, at least one wavelength emitted by the second laser source is congruent with at least one wavelength emitted by the first laser source. The refraction of light at the scanning optics, like the focusing of light by the focusing optics, is dependent on the wavelength. An excitation beam and an ablation beam with different wavelengths, which are guided into the scanning optics in a common beam path, can therefore be shifted chromatically with respect to one another and arrive at the sample with a spatial offset from one another. This chromatic shift is minimized if the wavelengths of both beams are identical.

Alternatively, the two beams can also have different wavelengths. They can then be brought together in particular via a dichromatic beam splitter with only low intensity losses.

In a further particularly advantageous embodiment of the invention, the polarization directions of the first laser source and the second laser source enclose an angle between 70 and 110 degrees. Both directions of polarization are preferably orthogonal to one another. The excitation beam and the ablation beam can then be brought together in particular via a polarization-maintaining beam splitter with only low intensity losses. The interaction of the two beams, in particular with biological samples that do not have a preferred crystalline direction, is generally independent of the direction of polarization. Furthermore, different directions of polarization of the excitation beam and the ablation beam when passing through the scanning optics and the focusing optics do not lead to an offset between the locations at which both beams arrive on the sample.

The first laser source and the second laser source are fed by a common continuous wave pump laser or a common pulsed pump laser source. In particular, the beam from the common continuous wave pump laser can be guided into an optical oscillator, which can in particular also be a fiber optic oscillator, and a beam splitter can be used to split the pulsed beam emitted by the optical oscillator into the excitation beam on the one hand and into the ablation beam on the other be provided. By sharing components for the excitation beam on the one hand and for the ablation beam on the other hand, costs, installation space and energy consumption can be saved. By using fiber-optic components, additional installation space can be saved. Furthermore, the adjustment is considerably simplified. Should there be two separate optical oscillators for the two due to the availability situation on the market energetically very different beams are cheaper than an oscillator equally suitable for both beams, it can also be advantageous to use two separate oscillators.

The excitation beam is passed through a spectral filter. For example, if the pulsed beam emitted by the optical oscillator has the very short pulse duration provided for the ablation beam, the spectral filter causes the pulses of the excitation beam to be significantly lengthened due to Heisenberg's uncertainty principle. At the same time, the spectral filter can also keep parts of the excitation beam away from the sample which are not suitable for changing the excitation states in the sample and thus only contribute to the heating of the sample.

The first laser source emits pulses of at least two different wavelengths. In particular, the first laser source can emit pulses of three different wavelengths. Such a laser source is particularly suitable for coherent anti-Stokes-Raman scattering. For this purpose, two of the emitted wavelengths advantageously have a difference that matches the excitation of at least one oscillation state in a molecule of the sample material. For example, a first emitted wavelength can be tunable in the range between 1025 nm and 1075 nm, and a second emitted wavelength can be tunable in the range between 800 nm and 1000 nm. The different wavelengths can be generated, for example, by a four-wave mixture of wavelengths that are symmetrically distributed around the wavelength of a pump laser used as an energy source. Ytterbium-doped fiber lasers, for example, are suitable for this. The four-wave mixing takes place in a photonic crystal fiber. The second laser source can, for example, also be such a fiber laser, the wavelength of which can be tuned, for example, in the range between 1030 nm and 1060 nm.

The detector is thus advantageously designed to detect light formed from the excitation beam by coherent Raman scattering, in particular stimulated Raman scattering (SRS) and / or by anti-Stokes Raman scattering.

1. a) Hierzu sind insbesondere Laserpulse im 10 ps-Pulsbereich optimal, da ein wichtiges Störsignal von SRS die Kreuzphasenmodulation ist, die proportional zur Zeitableitung der Feldstärke des Pulses ist. Daher ist das Störsignal für lange Pulse wesentlich geringer.
2. b) Weiterhin ist es vorteilhaft, einen stabilen Laser zu nutzen. Wenn der Modulationsübertrag auf den Yb-Faserlaser beobachtet wird, ist das besonders günstig, weil zur Erzeugung dieses Laserlichts keine nichtlinearen Effekte genutzt werden und daher das Rauschen minimal ist. Die Nutzung des Signals oder Idlers, der durch Vierwellenmischung erzeugt wird, wäre weniger günstig.
3. c) Es ist weiterhin vorteilhaft, nicht den Pumplaser zu detektieren, denn hier können Störsignale durch transiente Absorption auftreten. Bei SRS-Detektion auf dem Pumplaser wird SRL detektiert, was TA-Signalen äquivalent ist (stimulated Raman Loss). Diese Störsignale sind durch NIR-Laser und durch Detektion von Raman-Gain beim Stokes Laser statt Raman Loss vermeidbar.
4. d) Weiterhin ist es vorteilhaft, einen Laser mit (nahezu) fester Zentralfrequenz zu nutzen. Bei Verwendung des Yb-Lasers kann das SRS-Signal auch beim Durchstimmen der Raman-Resonanz detektiert werden, da die Yb-Wellenlänge um ca. 50 nm variiert wird, während die Signal-Wellenlängen in einem viel größeren Bereich von ca. 200 nm bzw. die Raman Resonanz von ca. 700-3300 cm^-1 durchgestimmt werden.

The laser source described above is particularly suitable for SRS imaging. To detect SRS, modulations of the pump laser are transmitted to the Stokes laser or vice versa through the non-linear Raman interaction in the sample. The modulation transfer is very small, typically below 10 ^-4 the laser intensity. Therefore, on the one hand, particularly low-noise lasers are required, and on the other hand, other sources of interference must be suppressed.

1. a) In particular, laser pulses in the 10 ps pulse range are optimal for this, since an important interference signal from SRS is the cross-phase modulation, which is proportional to the time derivative of the field strength of the pulse. The interference signal is therefore much lower for long pulses.
2. b) It is also advantageous to use a stable laser. If the modulation transfer to the Yb fiber laser is observed, this is particularly favorable because no non-linear effects are used to generate this laser light and the noise is therefore minimal. Using the signal or idler that is generated by four-wave mixing would be less beneficial.
3. c) It is also advantageous not to detect the pump laser, because interference signals due to transient absorption can occur here. With SRS detection on the pump laser, SRL is detected, which is equivalent to TA signals (stimulated Raman Loss). These interfering signals can be avoided by using NIR lasers and by detecting Raman gain with the Stokes laser instead of Raman loss.
4. d) It is also advantageous to use a laser with a (almost) fixed central frequency. When using the Yb laser, the SRS signal can also be detected when tuning the Raman resonance, since the Yb wavelength is varied by approx. 50 nm, while the signal wavelengths in a much larger range of approx. 200 nm or the Raman resonance of approx. 700-3300 cm ^-1 .

In a further particularly advantageous embodiment of the invention, the wavelength emitted by the first laser source and / or by the second laser source is between 750 nm and 3 μm, preferably between 750 nm and 2 μm and very particularly preferably between 750 nm and 1.5 µm. This wavelength range is particularly advantageous for the examination and modification of biological samples, since scattering losses in the tissue are minimized and the light can penetrate the sample to a depth of a few 100 µm.

The invention also relates to a method for operating a laser microscope, a pulsed excitation beam and a pulsed ablation beam being guided to a sample in the laser microscope, means for scanning the excitation beam and the ablation beam over the sample being provided. In the laser microscope, at least one detector is also provided for light which the sample emits due to a non-linear optical effect in response to the excitation beam.

According to the invention, the pulse duration of the ablation beam is selected between 35 fs and 300 fs, preferably between 100 fs and 300 fs.

It was recognized that a local ablation of the sample material can take place precisely with a pulse duration in this area without the sample otherwise being excessively heated. As explained above, the sample material is vaporized in that the instantaneous electrical field of the ablation pulse partially ionizes the electron shells of atoms of the sample material. This effect only occurs from a certain minimum field strength that is sufficient to overcome the binding energy of at least the external electrons. This minimum field strength corresponds to a minimum value for the instantaneous intensity of the ablation pulse (of the order of 10 ¹² -10 ¹⁴ W / cm ² ). During the course of the ablation pulse, the current intensity must reach this minimum value in a rising flank and drop again in a falling flank at the end of the pulse so that no other, in particular thermal, interaction of the ablation pulse with the sample material takes place on these flanks. The ablation pulse must therefore rise and fall again on a faster time scale than is necessary to excite oscillations or rotations in molecules of the sample material and in this way to couple heat into the sample material. If such an excitation of vibrations takes place, there is a high probability that the sample will be heated so much that it will be destroyed. The selective ablation is based on the fact that during the phase of the pulse in which a direct ionization of the electron shells of atoms takes place, at least one order of magnitude more energy is coupled into the sample than during the rising and falling edges of the pulse, during which the instantaneous intensity is not is sufficient for direct ionization. If sample material is locally ablated with pulses according to the invention, this can already be done, for example, with average powers of the ablation beam in the order of magnitude of 1 mW.

In a particularly advantageous embodiment of the invention, the pulse duration of the excitation beam for imaging is selected to be a factor between 10 and 1000 longer than the pulse duration of the ablation beam. This ensures that, on the one hand, the excitation beam does not ablate material through direct ionization and, on the other hand, that sufficient time is available to generate a certain excitation state in the sample through the interaction of the photons of the excitation beam with the sample. The qualitative difference between the effects of the excitation beam and the ablation beam is largely based on the different time scales and intensity scales on which these effects take place.

The pulse duration of the excitation beam is advantageously selected from a range between 1 ps and 100 ps, preferably between 5 ps and 40 ps and very particularly preferably between 10 ps and 20 ps. The adjustment of the excitation beam is easiest in this area. Furthermore, the ranges between 5 ps and 40 ps, or between 10 ps and 20 ps, are particularly advantageous when the excitation beam is guided through at least one optical fiber, for example when the first laser source is a fiber laser. Optical fibers typically have a dispersion around 10 ps per meter length for the two wavelengths required for imaging. From a pulse duration of around 10 ps, the dispersion in the optical fiber, combined with the dispersions in the scanning optics and in the focusing optics of the microscope, is low enough to perform the choreography of a spectroscopy with excitation by a pump pulse and query by an interrogation pulse ( Pump-probe spectroscopy) can no longer be decisively influenced when parts of the microscope optics such as objectives, scan lenses, tube lenses or condensers are changed. The range between 5 ps and 40 ps is also optimal with regard to the spectral resolution, which is directly linked to the pulse duration by the Heisenberg uncertainty principle, and at the same time guarantees the peak pulse powers in kW required for the non-linear processes with sufficient pulse repetition frequencies above 1 MHz for imaging Range for medium powers in the range of a few 10 mW.

The repetition rate of the pulses of the excitation beam is advantageously chosen between 1 MHz and 40 MHz, preferably between 1 MHz and 20 MHz. This area is an optimal compromise between the highest possible speed of image acquisition on the one hand and the lowest possible heating of the sample on the other. At least 8 million pulses per second are required for video frame rates. While during the ablation the energy coupled into the sample is essentially dissipated directly with the vaporized material and hardly leaves any heat in the sample, the excitation beam heats the sample afterwards According to its mean performance. At least one pulse of the excitation beam is required to record each image pixel. Depending on the signal-to-noise ratio of the effect selected for imaging, it can also be advantageous to provide several pulses of the excitation beam per image pixel in order to obtain better statistics.

The repetition rate of the pulses of the ablation beam is advantageously chosen between 100 kHz and 10 MHz, preferably between 100 kHz and 1 MHz. To ablate larger structures in a short time, the ablation beam can be configured, for example, in such a way that each pulse is effective in an area which comprises several image pixels, for example approximately 10 image pixels, of the image recorded with the excitation beam. The ablation beam can then be scanned over the sample more quickly, i.e. in a wider-meshed network of grid points. Ideally, the areas in which each ablation pulse removes material from the sample are seamlessly combined to form the structure to be ablated. During the ablation of the structure, for example, the size of the focus area of the ablation beam can also be varied in order to initially ablate large-area structures at high speed and then to rework fine structures with better accuracy.

The image field of the laser microscope can, for example, have an area of 1 mm ² . A spatial resolution of better than 1 μm laterally, that is to say along the surface of the sample, can typically be achieved. Axially, that is to say in the depth below the surface of the sample, a resolution of better than 5 µm can typically be achieved. The ablation beam can remove material with a resolution of typically 1-1000 µm ³ .

In a further particularly advantageous embodiment of the invention, the image obtained by scanning the excitation beam is evaluated using at least one multivariate classifier to determine whether the sample has a has a predetermined structure or property. For this purpose, for example, the European patent application going back to the applicant 15 200 864.5 disclosed classifier can be used.

A large canon of multivariate classifiers is available for many applications. The classifiers to be used can be selected from this canon, for example according to the required evaluation time, in order to be able to complete the evaluation within a specified time.

Furthermore, in biological in vivo applications, in which motion artifacts can occur, there may be specifications for a minimum speed of scanning in order to minimize the motion artifacts. An increased speed can then be accompanied by the fact that increased noise occurs. This noise can affect different multivariate classifiers to different degrees. In order to assess the reliability of the possible classifiers under the influence of image noise, the image obtained by scanning the excitation beam is changed in a further particularly advantageous embodiment of the invention by superimposing noise to form a test image. The reliability of the classifier is evaluated from the comparison of the results that the classifier delivers when applied to the image on the one hand and to the test image on the other hand.

For example, a classifier that changes its mind when there is even a slight amount of additional noise can be rated as less reliable than a classifier that only changes its mind when there is a lot of additional noise. The reliability determined in this way will typically depend on the type and strength of the noise contained in the image recorded by the laser microscope. The noise, in turn, depends on the speed of the image acquisition. Since the reliability can be checked quantitatively by adding additional noise, the user can of the laser microscope, choose an optimal compromise between the speed of image acquisition on the one hand and the usability of as many meaningful classifiers as possible on the other.

The method is carried out with a laser microscope of the invention. The laser microscope is specially designed to be operated with the method according to the invention, and vice versa.

One possible application of the laser microscope and the method according to the invention is the SRS or CARS-guided fs laser ablation of tissue for microsurgical operations based on the combination of multimodal non-linear microscopy (SRS, CARS, TPEF, SHG) and stimulated Raman Scattering microscopy (SRS microscopy) or coherent anti-Stokes' Raman scattering microscopy (CARS microscopy) of tissue for local diagnosis and characterization of the tissue with the targeted ablation of parts of the tissue by fs laser ablation. The main components in this application are the data acquisition and data processing, the combination of the imaging process with a tissue ablation process and the laser source used for this. The method enables the imaging and molecule-sensitive detection of target structures in tissue without the use of external marker substances and the subsequent precise ablation of the target structures. Tissue can be displayed non-destructively and three-dimensionally down to a depth of a few 100 µm and tissue can be removed in a targeted manner up to a few 100 µm below the surface, ie without creating an open wound. This can significantly reduce the risk of infection. The method is suitable for all body regions that are accessible to microscopes, eg skin, as well as for surgical interventions that are carried out with surgical microscopes, eg in the ear, nose and throat area, flexible or rigid imaging endoscopes.

The application combines marker-free molecular imaging for the localization of disease-related tissue anomalies with fs laser ablation for targeted ablation. This combines diagnostics and therapy in one device, which contributes to faster treatment. Furthermore, the fs laser ablation allows a much more precise removal of target structures and can also be used in endoscopes and microendoscopes. This makes the procedure particularly advantageous in the vicinity of physiologically important tissue structures, e.g. in the larynx near the vocal cords or in the brain.

The previous gold standard was based on the removal and histological processing of tissue biopsies to diagnose the disease and possibly subsequent further operations if the examination confirmed the suspicion of a serious disease. Thin tissue sections were made from the removed material and stained histologically, primarily using hematoxylin-eosin staining. The stained tissue section was assessed by a pathologist. This established process was time consuming and could take several days to complete. The accuracy of conventional surgical interventions and operations was limited to approx. 100 µm.

Due to the time-consuming sample processing, the success of the operation could not be checked during the operation, so that in some cases cost-intensive repeat operations were necessary. Usually, a lot of tissue was ablated, which increased the risk of infection and could damage important physiological structures.

Here the invention provides the combination of a marker-free imaging method, which can make molecules directly visible, with an optical method for tissue ablation. Using specific laser parameters, a high penetration depth of several 100 µm can be achieved for both imaging and laser ablation. The process is much faster and more accurate than conventional methods.

1. (i) Darstellung der Zielregion mit Hilfe der multimodalen nichtlinearen Mikroskopie, z.B. der kohärenten anti-Stokes Raman-Streuungsmikroskopie (bei einer oder mehreren Schwingungsfrequenzen) allein oder in Kombination mit der Zweiphotonenfluoreszenz und der zweiten Harmonischen
2. (ii) die multivariate Analyse der Bilddaten zur Erkennung der Zielregion für die Laserablation (basierend etwa auf der vorherigen europäischen Patentanmeldung 15 200 864.5 ) und
3. (iii) die lokale Abtragung von Zielgewebe und Gewebestrukturen in-vivo/excorpore-in-vivo/in-vitro/ex-vivo auch unterhalb einer intakten Gewebeschicht.

The new workflow using the invention includes:

1. (i) Representation of the target region using multimodal non-linear microscopy, e.g. coherent anti-Stokes Raman scattering microscopy (at one or more oscillation frequencies) alone or in combination with two-photon fluorescence and the second harmonic
2. (ii) the multivariate analysis of the image data to identify the target region for laser ablation (based on the previous European patent application, for example 15 200 864.5 ) and
3. (iii) the local ablation of target tissue and tissue structures in vivo / excorpore in vivo / in vitro / ex vivo also below an intact tissue layer.

The main innovation in this workflow is the combination of tissue measurement and display with multimodal non-linear microscopy (e.g. SRS, CARS, TPEF, SHG) with laser ablation for targeted tissue removal. The laser scanning microscope in combination with the compact laser source for SRS and CARS imaging and laser ablation is the most important instrument in this context. Both the combination of both processes and the construction of a compact, air-cooled high-power ablation laser are fundamentally new.

Furthermore, neither coherent Raman microscopy nor multimodal nonlinear microscopy has been used intraoperatively to date. Applications in animal experiments are also limited to imaging. A coupling with optical methods for targeted tissue removal and online monitoring of the progress of the operation is completely new.

• Untersuchung ausgedehnter intakter Gewebeproben, keine dünnen Schnellschnitte auf Objektträgern
• Detektion der Signale in Reflexion: da für ausgedehnte Gewebestrukturen keine Signaldetektion in Vorwärtsrichtung möglich ist, müssen die Signale in Rückwärtsrichtung erfasst werden
• Untersuchung von Gewebe in Echtzeit: da Bewegungsartefakte auftreten, werden die Untersuchungen mit höherer Geschwindigkeit im Vergleich zur histologischen Schnellschnittdiagnostik gemacht, dadurch ist mit höherem Rauschen zu rechnen und die automatisierte Datenanalyse wird auf wenige wichtige Parameter beschränkt
• Echtzeit-Analyse: Online-Datenverarbeitung direkt im Anschluss zur Datenaufnahme
• Anregung: NIR-Laser, 750-1500 nm - um eine hohe Eindringtiefe zu erreichen, wird eine langwellige Beleuchtung für die Bildgebung und fs-Ablation gewählt, um insbesondere Streuverluste im Gewebe zu minimieren.

Compared to the use of coherent Raman microscopy and multimodal nonlinear microscopy for rapid section diagnostics, for example according to the European patent application that goes back to the inventor 15 200 864.5 , the following differences are important:

• Examination of extensive intact tissue samples, no thin quick sections on microscope slides
• Detection of the signals in reflection: since signal detection in the forward direction is not possible for extensive tissue structures, the signals must be recorded in the reverse direction
• Examination of tissue in real time: since movement artifacts occur, the examinations are carried out at a higher speed compared to histological rapid section diagnostics, which means that higher noise is to be expected and the automated data analysis is limited to a few important parameters
• Real-time analysis: online data processing directly after the data acquisition
• Excitation: NIR laser, 750-1500 nm - in order to achieve a high penetration depth, long-wave illumination is selected for imaging and fs ablation, in particular to minimize scattering losses in the tissue.

• Pathologische Gewebestrukturen können in vivo detektiert und visualisiert werden, so dass eine Abgrenzung zum umgebenden gesunden Gewebe sichtbar gemacht werden kann.
• Pathologische Gewebestrukturen können gezielt mit µm-Ortsauflösung entfernt werden, auch in 3D und umgeben von gesundem Gewebe.
• Der neue Workflow ist auch für kritische Operationen an physiologisch wichtigen Strukturen geeignet, da mit extrem hoher Präzision operiert werden kann und Zielstrukturen kontrastreich dargestellt werden können. Auf Kontrastmittel wird verzichtet.
• Da die Untersuchung direkt im OP erfolgen kann, ermöglicht die Methode die Einsparung von Zeit und Geld, da auf eine Biopsieentnahme und Begutachtung verzichtet werden kann. Da der Erfolg der OP sofort geprüft werden kann, können Wiederholungs-OPs vermieden werden, was zu wesentlichen Kostenersparnissen in der operativen Patientenversorgung führen kann.

The new workflow, in which the laser microscope and the method according to the invention are the essential aids, has the following essential advantages:

• Pathological tissue structures can be detected and visualized in vivo so that a delimitation from the surrounding healthy tissue can be made visible.
• Pathological tissue structures can be specifically removed with µm spatial resolution, also in 3D and surrounded by healthy tissue.
• The new workflow is also suitable for critical operations on physiologically important structures, as operations are performed with extremely high precision can and target structures can be displayed with high contrast. Contrast media are not used.
• Since the examination can be carried out directly in the operating room, the method saves time and money, since a biopsy and assessment can be dispensed with. Since the success of the operation can be checked immediately, repeat operations can be avoided, which can lead to significant cost savings in operative patient care.

Special descriptive part

• Figur 1: Ausführungsbeispiel des Lasermikroskops 1 mit zwei separaten Laserquellen 10 und 20 für Anregung und Ablation, einem CARS-Detektor 61 in Transmission und einem CARS-Detektor 62 in Reflexion.
• Figur 2: Selektives Abtragen einer Ablagerung 82 von einer Arterienwand 81.
• Figur 3: Weiteres Ausführungsbeispiel des Lasermikroskops 1 mit einer gemeinsamen Laserquelle 10=20 für Anregung und Ablation sowie einem mehrstufigen CARS-Detektor 62 in Reflexion.
• Figur 4: Prüfung der Verlässlichkeit 31a-39a eines multivariaten Klassifizierers 31-39 durch Überlagerung mit Testrauschen 65.

The subject matter of the invention is explained below with reference to figures, without the subject matter of the invention being limited thereby. It is shown:

• Figure 1 : Embodiment of the laser microscope 1 with two separate laser sources 10 and 20 for excitation and ablation, a CARS detector 61 in transmission and a CARS detector 62 in reflection.
• Figure 2 : Selectively ablating a deposit 82 from an arterial wall 81.
• Figure 3 : Another embodiment of the laser microscope 1 with a common laser source 10 = 20 for excitation and ablation and a multi-stage CARS detector 62 in reflection.
• Figure 4 : Checking the reliability 31a-39a of a multivariate classifier 31-39 by superimposing test noise 65.

Figure 1 shows a first embodiment of a laser microscope 1 according to the invention. A first laser source 10 emits an excitation beam 11 composed of pulses 11a of a first wavelength and pulses 11b of a second wavelength, the difference between the wavelengths of the pulses 11a and 11b corresponding to the frequency of an oscillation in biological sample material 5a. A dichromatic beam splitter 91 directs the excitation beam 11 in the direction of the scanning optics 3. A second laser source 20 emits an ablation beam 21 made up of pulses from another Wavelength. The ablation beam 21 is also guided into the scanning optics 3 via a mirror 22 and the dichromatic beam splitter 91.

The excitation beam 11 and the ablation beam 21 are jointly guided from the scanning optics 3 into the focusing optics 4, which comprise a scanning and tube lens system 4a, a further dichromatic beam splitter 4b and an objective 4c. The beams 11 and 21 are focused together on the biological sample material 5a, which is applied as a thin layer on a slide 5b and together with the slide 5b forms the sample 5. The surface 55 of the sample 5 is approximately flat here.

Part of the light 11a, 11b, 21 radiated onto the sample 5 and the light 7 coherently Raman-scattered by the sample 5 is transmitted and reaches a first multimodal CARS detector 61. All of the light 7 passes through the CARS detector 61, 11a, 11b, 21 first a condenser 61a and is converted into a parallel beam path. A dichromatic beam splitter 61b separates the coherent anti-Stokes Raman scattered light 7a and guides it through a dielectric filter 63a, which in particular retains residual portions of all laser beams 11a, 11b and 21, to a first photomultiplier 61c. The light transmitted by the beam splitter 61b contains a further signal component 7b which is due to two-photon excitation fluorescence (TPEF), second harmonic generation (SHG) or another freely selectable optical effect. This signal component 7b is separated from the laser beams 11a, 11b, 21 with a further dielectric filter 63b and passed to a second photomultiplier 61d. If the dielectric filter 63b is removed, the laser light 11a, 11b, 21 can optionally be monitored for intensity fluctuations with the second photomultiplier 61d.

The dielectric filters 63a and 63b typically have an optical density of around 6 for the laser wavelengths used. They can be added as an option through another, in Figure 1 short-pass filter, not shown, which is arranged between the condenser 61a and the dichroic beam splitter 61b.

The light 7, 11a, 11b, 21 reflected by the sample passes the beam splitter 4b and reaches the second multimodal CARS detector 62. In the second CARS detector 62, the Raman-scattered portion 7a of the light is separated off with a dichromatic beam splitter 62x and passed to a photomultiplier 62y via a dielectric filter 63c, which in particular retains residual components of all laser beams 11a, 11b and 21. Analogously to the first CARS detector 61 operated in the transmission configuration, the light transmitted by the beam splitter 62x contains the signal component 7b. This signal component 7b is separated from the laser beams 11a, 11b, 21 with a further dielectric filter 63d and reaches a photodiode 62z. If the dielectric filter 63d is removed, the photodiode 62z can be used to monitor the laser beams 11a, 11b and 21 for intensity fluctuations or, for example, to normalize the Raman spectra to the total intensity. Due to its larger dynamic range, the photodiode 62z is more suitable for this than a photomultiplier 61a, 61b, 62y. In addition, the photodiode 62z can be used with a suitable filter for a laser wavelength for SRS detection in combination with a lock-in amplifier or a tuned amplifier .

Conventional photomultipliers with secondary electron multipliers can be used as photomultipliers 61a, 61b and 62y. Alternatively, hybrid detectors can be used instead. In hybrid detectors of this type, primary electrons are generated in a cathode, which can consist, for example, of gallium arsenide phosphide. The primary electrons are then accelerated to a material that releases secondary electrons using a voltage that is significantly higher than that of conventional photomultipliers (approx. 5-10 kV). The secondary electrons are then directed to a diode and converted into a current pulse by this diode.

Figure 2 illustrates the selective ablation with the ablation beam 21. Figure 2a shows schematically a first, with the laser microscope 1 according to Figure 1 Recorded image 64 of a thin section of an arterial wall 81. In the interior 83 of the artery, deposits 82 have accumulated on the inside of the arterial wall 81.

Figure 2b shows schematically a further image 64 of the same image field after the selective removal of the deposits 82 with the ablation beam 21. The arterial wall 81 itself is undamaged.

That of the schematic Figure 2 The underlying real CARS recordings for a vibration resonance of 2850 cm ^-1 are to be understood as a "proof of concept" for the basic feasibility of a selective tissue removal. In real in-vivo applications, the sample, including the structures to be ablated, is not present as a thin section, but as a three-dimensional object.

A laser microscope 1 adapted to real in vivo use is shown in FIG Figure 3a outlined. In contrast to Figure 1 Here a single laser 10 = 20 is the common source for the excitation beam 11 and the ablation beam 21. This common laser 10 = 20 is significantly more compact than the arrangement of two separate lasers 10, 20 according to FIG Figure 1 .

Still in contrast to Figure 1 the ablation beam 21 has a wavelength which is also contained in the excitation beam 11. However, the direction of polarization of the ablation beam 21 is orthogonal to the direction of polarization of the excitation beam 11. Therefore, the excitation beam 11 and the ablation beam 21 are combined via a polarization-maintaining beam splitter 92.

The excitation beam 11 and the ablation beam 21 are analogous to FIG Figure 1 guided via the common scanning optics 3 and the common focusing optics 4 to the sample 5. The sample 5 here is a three-dimensional tissue object 5d on which a structure 5c to be ablated is indicated. Accordingly, the surface 55 of the sample 5 is also three-dimensional.

In contrast to Figure 1 sample 5 is not transparent. Therefore, measurements can only be made in reflection. The light 11a, 11b, 21 reflected by the sample 5 passes, together with the signal light 7 generated by the sample 5, the beam splitter 4b of the focusing optics 4 and arrives at the single multimodal CARS detector 62. In this CARS detector 62, the various signals , ie the Raman-scattered light 7a, SHG signals 7b, TPEF signals 7c, a further signal component 7d, as well as the laser light 11a, 11b, 21 with several cascaded dichromatic beam splitters 62a, 62b and 62d as well as suitable dielectric filters 63a, 63b, 63c and 63d separated.

The first dichromatic beam splitter 62a splits off a first wavelength component 7a of the signal light 7 and guides it via the dielectric filter 63a to the photomultiplier 62f. The remaining wavelength components 7b and 7c, e.g. TPEF and SHG, the reflected excitation light 11a, 11b and the reflected ablation beam 21 pass through the first dichromatic beam splitter 62a unhindered in the forward direction (in Figure 3a vertically up).

The second dichromatic beam splitter 62b splits a second wavelength component 7b and a third wavelength component 7c of the signal light 7. These two wavelength components 7b and 7c are then separated from one another in a third dichromatic beam splitter 62d and fed to the photomultipliers 62e and 62c via dielectric filters 63b and 63c, which only allow wavelength components 7b and 7c to pass and hide further spectral components. The reflected excitation light 11a, 11b and the reflected ablation beam 21 pass the second dichromatic beam splitter 62b together with a further signal component 7d again unhindered. The dielectric filter 63d masks out the laser light 11a, 11b, 21 so that only the signal component 7d reaches the photodiode 62q. Optionally, the dielectric filter 63d can be removed so that the photodiode 62q can be used to measure the intensity of the laser light 11a, 11b, 21. This intensity can then be analogous to Figure 1 can be used to control and normalize the Raman and other non-linear signals to the total intensity.

The advantage of the CARS detector 62 according to FIG Figure 3 is that this means that four wavelength components 7a, 7b, 7c and 7d of the signal light 7, for example CARS, SHG, TPEF and another freely selectable signal component, can be registered at the same time. These four wavelength components 7a, 7b, 7c and 7d can genuinely be generated simultaneously by the sample. However, they can also be generated one after the other, for example by tuning the wavelengths of the pulses 11a and 11b that form the excitation beam 11.

Figure 3b shows the internal structure of the common laser 10 = 20. This laser 10 = 20 is characterized in that most of the optical components are used both for the excitation beam 11 and for the ablation beam 21. The beam from a common continuous wave pump laser 15 is guided into an optical oscillator 16 and there is converted into pulses with the pulse duration suitable for the ablation beam 21. The beam 17 formed from these pulses is guided from the optical oscillator 16 to a beam splitter 18.

The beam splitter 18 lets the ablation beam 21 in the forward direction (in Figure 3b vertically upwards). The ablation beam 21 is amplified with an amplifier 18b and finally emerges from the laser 10 = 20.

The excitation beam 11 is directed laterally to a mirror 18a and from there to a spectral filter 19. Due to the Heisenberg uncertainty principle, the spectral filter 19 leads to the pulses of the excitation beam 11 being significantly lengthened. The excitation beam 11 is first amplified in an amplifier 19a.

In a photonic crystal fiber 19b, two further wavelengths are now generated from the excitation beam 11, which has essentially only one frequency ω ₀ behind the spectral filter 19, by four-wave mixing, signal and idler. Through the non-linear process of four-wave mixing, two photons of the excitation beam with frequency ω ₀ generate a pair of a signal photon with frequency ω ₀ + Δω and an idler photon with frequency ω ₀ -Δω. The photonic crystal fiber 19b is microstructured in such a way that, in spite of the dispersion in the crystal fiber 19b, the energy and momentum are conserved.

In the photonic crystal fiber 19b, broadband photon pairs ω ₀ ± Δω with many frequency shifts Δω are generated. So that exactly one frequency shift Δω is preferred and thus exit the laser 10 = 20 pulses 11a, 11b with exactly two frequencies (and thus two wavelengths), part of the light emerging from the photonic crystal fiber 19b is transferred to the photonic via a resonant cavity 19c Crystal fiber 19b fed back. The cavity 19c is only ever resonant at one frequency, that is to say either at the frequency ω ₀ + Δω or at the frequency ω ₀ -Δω. By thus defining the frequency shift Δω, both frequencies of the photon _{pair ω 0} ± Δω are determined, which should preferably be formed. The frequency shift Δω can be tuned through the resonance of the cavity 19c.

Figure 4 shows schematically how classifiers 31-39 can be examined to determine whether they are used for recognizing a given structure or property 41-49 in a concrete image 64, which is afflicted with noise 64a and which was recorded with the laser microscope 1. The image 64 is changed to a test image 66 with additional test noise 65. The classifier 31-39 is now applied on the one hand to the original image 64 and delivers a result 67 which includes the determination of whether the structure or property 41-49 is present on or in the sample 5 according to the original image 64. The classifier 31-39 is applied in parallel to the test image 66 and delivers a result 68. The two results are compared in block 69, and the reliability 31a-39a of the classifier 31-39 is evaluated from this comparison. This reliability 31a-39a can in particular depend on the strength of the additional noise 65a from which the classifier 31-39 changes its mind. If even a slight additional noise is sufficient for this, the conclusion can be drawn from this that the original noise 64a in the image 64 may have already falsified the result 67 supplied by the classifier 31-39. If, on the other hand, the opinion of the classifier 31-39 does not change even with strong noise, it can be concluded from this that the classifier is particularly resistant to noise and is therefore particularly reliable.

The disclosure also includes the following example after that described above:
A laser microscope comprising at least one first laser source that emits at least one, in particular pulsed, excitation beam, scanning optics that are designed to raster the excitation beam over the surface of a sample, focusing optics that are designed to focus the excitation beam on the sample, and At least one detector for light which the sample emits due to an optical effect in response to the excitation beam, a second laser source for a pulsed ablation beam being provided for the purpose of local ablation of the material of the sample, and the ablation beam via the scanning optics and the focusing optics to the sample is led.

List of reference symbols

1

Laser microscope

10

first laser source for excitation beam 11

11

Excitation beam

11a, 11b

Pulses of the beam 11 with different wavelengths

15th

common continuous wave pump laser for lasers 10, 20

16

optical oscillator

17th

Optical oscillator beam 16

18th

Beam splitter for splitting beam 17 into beams 11, 21

18a

Mirror for deflecting the excitation beam 11

18b

Ablation beam amplifier 21

19th

spectral filter for excitation beam 11

19a

Amplifier for excitation beam 11

19b

photonic crystal fiber for the formation of photon pairs ω ₀ ± Δω

19c

resonant cavity to select a frequency shift Δω

20th

second laser source for ablation beam 21

21

Ablation beam

22nd

Ablation beam mirror 21

3

Scanning optics or means for rasterizing

31-39

multivariate classifiers

31a-39a

Reliability of the multivariate classifiers 31-39

4th

Focusing optics

4a

Scan and tube lens system

4b

Beam splitter in focusing optics 4

4c

lens

41-49

Properties to which classifiers 31-39 are sensitive

5

sample

5a

thin section of biological material

5b

Microscope slide

5c

Structure to be ablated on sample 5

5d

three-dimensional object as a sample 5

55

Surface of the sample 5

61

CARS detector in transmission

61a

Condenser

61b

Beam splitter in CARS detector 61

61c, 61d

Photomultiplier in CARS detector 61

62

CARS detector in reflection

62a

first dichromatic beam splitter for splitting off 7a

62b

second dichromatic beam splitter for splitting off 7b, 7c

62c

Photomultiplier for wavelength component 7c

62d

third dichromatic beam splitter, separates 7b from 7c

62e

Photomultiplier for wavelength component 7b

62f

Photomultiplier for wavelength component 7a

62q

Photodiode for wavelength component 7d

62x

Beam splitter in simple CARS detector 62

62y

Photomultiplier in a simple CARS detector 62

62z

Photodiode in simple CARS detector 62

63a-63d

dielectric filters

64

Image recorded by the laser microscope 1

64a

Noise in picture 64

65

Test noise

66

Test image, generated from image 64 and test noise 65

67

Result of classifier 31-39 at picture 64

68

Result from classifier 31-39 on test image 66

69

Comparison of the results 67, 68

7th

response generated by sample 5

7a-7d

Wavelength components of the answer 7

81

Arterial wall

82

Deposits on arterial wall 81

83

Interior of the artery bounded by arterial wall 81

91

dichromatic beam splitter to merge 11, 21

92

polarization-maintaining beam splitter for merging 11, 21

ω0

Frequency of the excitation beam 11 behind the spectral filter 19

Δω

Frequency shift in photonic crystal fiber 19b

Claims (12)

1. Laser microscope (1), comprising a common laser (10, 20) emitting at least one excitation beam (11), a scan optics (3) configured for scanning the excitation beam (11) over the surface (55) of a sample (5), a focusing optics (4) configured for focusing the excitation beam (11) onto the sample (5), and at least one detector (61, 62) for light (7, 7a-7d) emitted by the sample (5) due to an optical effect in response to the excitation beam (11), wherein the common laser emits a pulsed ablation beam (21) for the purposes of local ablation of the material of the sample (5), wherein the ablation beam (21) is guided via the scan optics (3) and the focusing optics (4) to the sample (5), wherein the common laser (10, 20) comprises a common continuous-wave pump laser (15), wherein the excitation beam comprises pulses (11a, 11b) of at least two different wavelengths, wherein the common pulsed laser comprises an optical oscillator (16), a spectral filter (19) and a beam splitter (18), wherein the beam is guided from the common continuous-wave pump laser (15) into the optical oscillator (16) and the beam splitter (18) splits the pulsed beam (17) emitted by the optical oscillator (16) into the excitation beam (11) on the one hand and into the ablation beam (21) on the other hand, wherein the excitation beam (11) is guided through the spectral filter (19) and the at least two different wavelengths of the excitation beam (11) are generated by a four-wave mixture of the excitation light guided through the spectral filter (19).
2. Laser microscope (1) according to Claim 1, characterized in that at least one wavelength of the ablation beam (21) is congruent with at least one wavelength of the excitation beam (11).
3. Laser microscope according to one of Claims 1 to 2, characterized in that the detector (61, 62) is configured to detect light formed by coherent Raman radiation, in particular by CARS and SRS, from the excitation beam (11) .
4. Laser microscope (1) according to one of Claims 1 to 3, characterized in that the wavelength of the excitation beam (11) and/or the wavelength of the ablation beam (21) is between 750 nm and 3 µm, preferably between 750 nm and 2 µm and most preferably between 750 nm and 1.5 µm.
5. Laser microscope (1) according to one of Claims 1 to 4, characterized in that the detector (61, 62) is sensitive to light (7, 7a-7d) emitted by the sample (5) due to a non-linear optical effect in response to the excitation beam (11).
6. Method for operating a laser microscope (1), wherein in the laser microscope (1) an excitation beam (11) and a pulsed ablation beam (21), which are emitted by a common laser (10, 20), are guided to a sample (5), wherein the common pulsed laser (10, 20) comprises an optical oscillator (16), a spectral filter (19), a beam splitter (18) and a common continuous-wave pump laser (15), wherein the beam is guided from the common continuous-wave pump laser (15) into the optical oscillator (16) and the beam splitter (18) splits the pulsed beam (17) emitted by the optical oscillator (16) into the excitation beam (11) on the one hand and into the ablation beam (21) on the other hand, wherein the excitation beam (11) comprises pulses (11a, 11b) of at least two different wavelengths, wherein the excitation beam (11) is guided through the spectral filter (19) and the at least two different wavelengths of the excitation beam (11) are generated by a four-wave mixture of the excitation light guided through the spectral filter (19), wherein means (3) each for scanning the excitation beam (11) and the ablation beam (21) over the sample are provided and wherein at least one detector (61, 62) is provided for light emitted by the sample (5) due to a non-linear optical effect in response to the excitation beam (11), wherein the pulsed ablation beam (21) is configured to partially ionize the electron shells of atoms of the material of the sample (5) by its instantaneous electric field, wherein the pulse duration of the ablation beam (21) is selected between 35 fs and 300 fs, preferably between 100 fs and 300 fs.
7. Method according to Claim 6, characterized in that the pulse duration of the excitation beam (11) is selected to be longer by a factor of between 10 and 1000 than the pulse duration of the ablation beam (21).
8. Method according to one of Claims 6 to 7, characterized in that the pulse duration of the excitation beam (11) is selected from a range between 1 ps and 100 ps, preferably between 5 ps and 40 ps and most preferably between 10 ps and 20 ps.
9. Method according to one of Claims 6 to 8, characterized in that the repetition rate of the pulses of the excitation beam (11) is selected between 1 MHz and 40 MHz, preferably between 1 MHz and 20 MHz.
10. Method according to one of Claims 6 to 9, characterized in that the repetition rate of the pulses of the ablation beam (21) is selected between 100 kHz to 10 MHz, preferably between 100 kHz and 1 MHz.
11. Method according to one of Claims 6 to 10, characterized in that the image (64) obtained by scanning of the excitation beam (11) is then evaluated by application of at least one multivariate classifier (31-39) as to whether the sample (5) comprises a specified structure or property (41-49).
12. Method according to Claim 11, characterized in that the image (64) is changed by superposition of noise (65) to a test image (66) and the reliability (31a-39a) of the classifier (31-39) is evaluated from the comparison (69) of the results that are provided by the classifier (31-39) when applied to the image (64) on the one hand (67) and when applied to the test image (66) on the other hand (68) .

Images